System and method for noninvasively monitoring conditions of a subject

ABSTRACT

A method and system are presented for use in determining one or more parameters of a subject. A region of interest of the subject is irradiated with acoustic tagging radiation, which comprises at least one acoustic tagging beam. At least a portion of the region of interest is irradiated with at least one electromagnetic beam of a predetermined frequency range. Electromagnetic radiation response of the at least portion of the region of interest is detected and measured data indicative thereof is generated. The detected response comprises electromagnetic radiation tagged by the acoustic radiation. This enables processing of the measured data indicative of the detected electromagnetic radiation response to determine at least one parameter of the subject in a region corresponding to the locations in the medium at which the electromagnetic radiation has been tagged by the acoustic radiation, and outputting data indicative of the at least one determined parameter.

FIELD OF THE INVENTION

This invention relates to a method and system for monitoring a subject'scondition, based on scattered light distribution through turbid media.The invention is particularly useful in medical applications.

BACKGROUND OF THE INVENTION

Non invasive monitoring and imaging using non-ionizing radiation, allowsmedical professionals to diagnose and monitor a patient without invasivesurgeries, or even without drawing blood. Pulse oximetry is one suchrevolutionizing technology, where non invasive monitoring of bloodoxygenation using light has replaced blood gas analysis. Thus, pulseoximetry has become a gold standard monitor in every clinical setting,and has saved millions of lives.

During non-invasive monitoring, the concentration of certainchromophores (such as oxygenated and deoxygenated hemoglobin inoximetry) is calculated by detecting light that escapes the tissue,determining the optical properties of the tissue, and deriving therefromthe concentrations of the chromophores. Providing the tissue ishomogenous, simple models allow for the calculation of theseconcentrations. However, as biological tissue is a complex scatteringmedium, measuring the local optical properties becomes a challengingtask.

As light is highly scattered while propagating through turbid media suchas biological tissue, photons that escape the tissue and reach adetector do not provide information about the path that they followed asthey propagated through the medium. To acquire information about theoptical properties of the tissue in the photons' path, several methodsand algorithms have been developed. Such methods includefrequency-domain spectroscopy, and photoacoustic spectroscopy [D MHueber et al Phys. Med. Biol. 46 (2001) 41-62].

SUMMARY OF THE INVENTION

The present invention utilizes the principles of ultrasound tagging oflight. More specifically, the tagging of light by acoustic radiation isused to determine the optical response of a region of interest. Theinvention may be used, for example, to significantly improve oximetryand pulse oximetry based measurements.

According to the invention, a region of interest in a subject (e.g.human body) is illuminated with at least one wavelength of light, and isirradiated with acoustic radiation (preferably ultrasound) such that theacoustic radiation overlaps the illuminated region in at least a part ofthe region of interest during the duration of illumination and/ordetection of the illuminating light (this overlapping volume is termed“tagged volume”). This acoustic radiation is termed acoustic taggingradiation. Light scattered from the subject's body and including photonsthat are tagged by the acoustic radiation and those that are not, isappropriately detected.

It is a common goal of any optical measurement technique to be capableof providing a high resolution measurement of the local lightdistribution with an improved signal to noise ratio (SNR). The presentinvention addresses this problem by providing a novel method and systembased on the principles of acoustic tagging of light, where the acousticradiation is appropriately modulated (coded) to provide high-resolutionand high-SNR measurement results.

The main idea of the present invention is based on the followingunderstanding: The effect termed “Ultrasound Tagging of Light” (UTL) isbased on the interaction of acoustic waves with the same tissue volumethat is being probed by light. This interaction causes the light wave tobe modulated, or tagged, with the characteristics of the acoustic wave(i.e. frequency, phase). As the propagation of acoustic waves in tissueis relatively slow (about 1500 m/sec in soft tissue), the location ofthe interaction of light with the acoustic radiation can be determined.The efficiency and power of the interaction of the acoustic waves withthe medium affects the spatial and temporal resolution and the SNR ofthe measurement. There are three possible modalities for the generationof acoustic waves, a continuous wave (CW), a short burst of waves (SB),and a pulse. Operation with continuous waves produces a higher SNR. Whena continuous acoustic wave (at a predetermined frequency range)interacts with light, and light is collected throughout the fullpropagation of the acoustic waves, a higher acoustic energy is availablefor the interaction, thereby increasing the signal. In addition, thespectral bandwidth of the continuous acoustic wave can be very narrow,thus reducing noise bandwidth. Thereby the SNR is greatly improved.However, the spatial resolution of a measurement produced withcontinuous acoustic waves is not as high as a measurement produced withshort bursts or pulses of acoustic waves. This reduced spatialresolution is particularly limiting when the measurement geometry callsfor propagation of acoustic waves essentially parallel to the directionof light propagation. As for the use of short bursts of waves andpulses, this provides better spatial resolution, but the acoustic energyof the interaction is lower and the bandwidth is wider as compared tothose of a continuous wave mode, resulting in reduced SNR.

There is accordingly a need in the art for a measurement technique whichcan achieve both high spatial resolution and high SNR. The presentinvention solves this problem by utilizing generation of continuousacoustic waves (and therefore improving the SNR), where the continuousacoustic wave is a modulated (coded) signal characterized by a narrowautocorrelation function, thereby improving the spatial resolution.

The expression “narrow autocorrelation function” refers toautocorrelation which is negligible for any delay time larger than thedetermined time resolution of the system. The latter may for example bedetermined as the time resolution of detection of the electromagneticradiation response, or as the temporal bandwidth of the acousticexcitation of the ultrasound transducer, or as the required spatialresolution divided by the speed of sound in the media.

In some embodiments of the invention, a pseudo random sequence, orspecially designed sequences such as Barker codes, or Golay codes (usedin radar technology) can be used. A combination of several sucharbitrary signals (having different phases and/or amplitudes) can beused interchangeably. According to one specific but not limitingexample, the modulated signal may be a non-periodic time function withpredefined time intervals between such non-periodic occurrences.

In some embodiments of the present invention, the coding comprises aseries of short pulses with high amplitude, that are separated byperiods of low amplitude (or even zero amplitude). The duration of thehigh amplitude pulses depends on the required time resolution of thesystem. The separation duration between two consecutive pulses isdetermined such that the phase of light propagating through the mediaduring the second pulse is independent of the phase of light during theprevious pulse of acoustic radiation. In addition, the consecutive highamplitude pulses may differ in frequency or may also be chirped.

The present invention thus provides for a 3D mapping of the lightdistribution in a turbid medium, obtaining a non invasive means forcollecting data about the structure and composition of the turbidmedium. The use of a continuous acoustic signal utilizes the acousticand light energy more efficiently, and lower acoustic and opticalsignals can be used while maintaining the desired SNR. Thus, the lightlevels and acoustic levels introduced into the subject are safer.

According to one broad aspect of the invention, a method is provided foruse in to determining one or more parameters of a subject's tissue, themethod comprising:

(a) irradiating a region of interest of the subject with acoustictagging radiation, the acoustic tagging radiation comprising at leastone acoustic tagging beam being a coded continuous acoustic wave in theform of a predetermined function of at least one parameter of theacoustic radiation varying over time during a measurement time interval,said predetermined time function having a narrow autocorrelation;

(b) irradiating at least a portion of the region of interest with atleast one electromagnetic beam of a predetermined frequency range;

(c) detecting an electromagnetic radiation response of said at leastportion of the region of interest and generating data indicativethereof, said response comprising electromagnetic radiation tagged bythe acoustic radiation, thereby enabling processing said data indicativeof the detected electromagnetic radiation response, to determine atleast one parameter of the subject's tissue in a region corresponding tothe locations in the medium at which the electromagnetic radiation hasbeen tagged by the acoustic radiation, and output data indicative of theat least one determined parameter.

According to another broad aspect of the invention, there is provided asystem for use in determining one or more parameters of a subject, thesystem comprising:

an acoustic unit configured and operable for irradiating a region ofinterest with acoustic tagging radiation comprising at least oneacoustic tagging beam being a coded continuous acoustic wave in the formof a predetermined function of at least one parameter of the acousticradiation varying over time during a predetermined time interval usedfor measurements, said predetermined function having narrowautocorrelation; and

an optical unit configured and operable for irradiating at least aportion of the region of interest with at least one electromagnetic beamof a predetermined frequency range, detecting an electromagneticradiation response of said at least portion of the region of interestand generating data indicative thereof, said response comprisingelectromagnetic radiation tagged by the acoustic radiation, said databeing indicative of the at least one parameter of the subject in aregion corresponding to the locations in the medium at which theelectromagnetic radiation has been tagged by the acoustic radiation.

The generation of such a coded acoustic wave can be implemented asfollows:

An arbitrary sequence can be produced and stored, the arbitrary sequenceactivating an arbitrary waveform generator. The latter (or anappropriate arbitrary switch) thus generates an arbitrary sequence ofelectronic signals which corresponds to the stored arbitrary sequence.Such an electronic signal in the form of an arbitrary sequence presentsa modulating or coding signal for operating an acoustic transducer. Theoutput of the acoustic transducer thus generated is a correspondingmodulated acoustic wave. The arbitrary sequence used for generation of amodulating signal can incorporate modulations of the original signal infrequency and/or phase and/or amplitude and/or any other parametricdomain. The modulated signal should have a narrow autocorrelation thatdefines the time resolution of the detection. As indicated above, thismay be a pseudo random sequence, or specially designed sequences (suchas Barker codes, or Golay codes used in radar technology), or acombination of several such arbitrary signals having different phasesand/or amplitudes used interchangeably.

The detection of the light response of the medium is implemented usingone or more appropriate photodetectors, each for receiving lightreturned (scattered) from the medium and generating an electronic outputcorresponding to the detected light intensity. Light collected by thedetector includes both tagged and untagged photons. The electronicoutput signal of the detector is processed by correlating it with theoriginal modulated signal (stored arbitrary sequence).

According to some embodiments of the present invention, the correlationis done using a cross correlation function to determine the opticalproperties of the medium at different depths. To this end, the crosscorrelation is determined for different time delays from the onset ofthe acoustic wave. At each delay, the cross correlation represents theintensity of tagged light corresponding to a distance from the acoustictransducer (e.g. depth in the subject) equal to the product of the speedof sound in the subject's tissue and the delay time. Since the processof acoustic tagging of light does not have a constant phase relationwith the acoustic tagging signal, a phase matching mechanism ispreferably added to the cross correlating algorithm. The amplitude ofthe cross correlation at each delay is assumed to correspond to afunction of the light distribution at the corresponding depth and thepressure amplitude of the acoustic wave at that depth. For example, thisfunction corresponds to the product of the two parameters. The lightdistribution can be determined by eliminating the contribution of theacoustic wave distribution to the amplitude of the cross correlation. Byfitting the light distribution to an expected distribution (for example,an exponential attenuation), the optical properties of the layer wherethe amplitude of the cross correlation is measured are determined.

In some embodiments of the present invention, multiple light sourcesand/or detectors and/or acoustic sources may be used. Suchconfigurations improve the spatial resolution of the measurements andenable the mapping of a larger volume of the medium. For the purposes ofthe present invention, when multiple acoustic sources (i.e. multipleacoustic waves) are used, all acoustic sources can use either differentfrequency ranges or the same frequency range, as long as the modulatingsequences of the different acoustic sources have zero or near zero crosscorrelation. When the electromagnetic response signal is detected anddecoded, each acoustic beam tagging effect can be estimated separatelyby correlating the respective received signal with the originalmodulating sequence for this acoustic source. The contribution of otheracoustic sources to such a correlation is negligible given the zero ornear zero cross correlation between the sequences, as will be describedbelow.

When multiple acoustic sources are used, they may be arranged andoperated such that acoustic radiations produced by these sourcesinterfere in at least a volume part of the region of interest. By this,the acoustic power in that volume can be enhanced or nullified accordingto the desired application. In this case, the different acoustic signalsgenerated by different acoustic sources are selected such as to providenon-zero cross correlation thereof at a predetermined delay. Thus, inthe region of interest, the overall acoustic radiation is a combinationof several acoustic signals.

The present invention can be used for various applications, includingmedical and non-medical ones. Considering the medical applications, thepresent invention can be used for example for determining oxygensaturation in blood and/or tissue, as well as determining concentrationof substance(s) in blood and/or tissue such as hemoglobin, glucose, etc.As an example, the invention is used in the determination of oxygensaturation of the tissue layers, and is therefore described below withrespect to this specific application, but it should be understood thatthe invention is not limited to this specific application.

BRIEF DESCRIPTION OF THE DRAWINGS

In order to understand the invention and to see how it may be carriedout in to practice, a preferred embodiment will now be described, by wayof non-limiting example only, with reference to the accompanyingdrawings, in which:

FIG. 1 is a schematic illustration of a measurement system according toan embodiment of the present invention;

FIG. 2A is a block diagram of an example of a control unit for use inthe system of the present invention;

FIG. 2B is a flow chart of an example of a method of the presentinvention;

FIGS. 3A to 3C exemplify generation of a phase coded continuous acousticsignal, where FIG. 3A shows a segment of an exemplary signal, FIG. 3Bshows auto correlation of said signal, and FIG. 3C shows the correlationC(τ₀) for time delay τ₀=10⁻⁵ seconds;

FIGS. 4A to 4C similarly show an example of generation of a frequencycoded continuous acoustic signal;

FIG. 5 shows an example of the cross correlation of a synthetic tissuemodel using phase modulated continuous acoustic signal for threedifferent light wavelengths;

FIGS. 6A-6C and 7A-7B show examples of different configurations of aprobe device according to the invention; and

FIGS. 8A and 8B show an example of transducer's assemblies includinglight guides.

DETAILED DESCRIPTION OF EXEMPLARY EMBODIMENTS

Reference is made to FIG. 1 illustrating schematically a specific butnot limiting example of a measurement system, generally designated 100,configured and operable according to the invention for non-invasivedetermination of one or more parameters (properties of tissuecomponents) of a subject, particularly a human or animal body. Theparameter(s) to be determined may include oxygen saturation level, orvalues/levels of various other parameters such as the concentration ofan analyte in the patient's blood, or the perfusion of ananalyte/metabolite in tissues. The values of these parameters arederived from the light distribution in a region of interest 200 as willbe described below.

System 100 includes such main constructional parts as a measurement unit101 and a control unit 120. Measurement unit 101 includes an optical orelectromagnetic unit (module) 101C and an acoustic unit (module) 110.Optical module 101C includes an illumination assembly 101A and a lightdetection assembly 102A, and acoustic module 110 is configured as anacoustic transducer arrangement including one or more acoustictransducers. Control unit 120 is configured to control the operation ofmeasurement unit 101, and to process, analyze measured data generated bymeasurement unit 101 (its detection assembly), and display the resultsof the analysis.

Illumination assembly 101A includes one or more illumination sourcesassociated with one or more different locations with respect to theregion of interest. Similarly, detection assembly 102A includes one ormore detector units associated with one or more different detectinglocations. It should be noted that the illumination source includes oneor more lighting elements each formed, for example, by a light emitterand possibly also a light guiding unit (e.g., an optical fiber or fiberbundle). For example, a probe part of the measurement unit by which itis to be brought to the body part under measurements may carry the lightemitter itself, or may carry a distal end of a light guiding unit whichby its opposite end is connected to an external light emitter. Thedetector unit includes one or more light detecting elements each formedby a light sensor and possibly also a light guiding unit (e.g., opticalfiber or fiber bundle); the probe part by which the measurement unit isto be brought to the body part may carry the light sensor or a distalend of the light guide which by its opposite end is coupled to anexternal light sensor.

The lighting element(s) and/or detecting element(s) may be incorporatedwithin the acoustic transducer arrangement as will be described furtherbelow with reference to FIGS. 6A-6C and 7A-7B. The lighting element(s)and detecting element(s) may be incorporated in one unit and theacoustic transducer arrangement placed to the side of (and not inbetween) the detector element(s) and the lighting element(s).

In the example of FIG. 1, illumination assembly 101A includes a singleillumination unit, and light detection assembly 102A includes a singledetector unit. It should be understood that this does not necessarilysignify the use of a single lighting element and/or a single detectingelement. Such a single illumination unit, as well as a single detectionunit, may include an array of lighting elements and an array ofdetecting elements, such that all the lighting elements of the sameillumination unit are to associated with the same location with respectto the region of interest, and similarly all the detecting elements ofthe same detection unit are associated with the same location relativeto the region of interest. In some other embodiments of the invention,the illumination assembly includes more than one illumination unitand/or more than one detection unit, as will be described below.

Optical module 101C and acoustic module 110 are connected to controlunit 120, e.g., by cables 105, 106 and 107 as shown in FIG. 1, or usingwireless signal transmission (e.g. IR, RF or acoustic signaltransmission) as the case may be.

Control unit 120 is typically a computerized system including inter aliaa power supply unit (not shown); a control panel with input/outputfunctions (not shown); a data presentation utility (e.g. display) 120A;a memory utility 120B; and a data processing and analyzing utility (e.g.CPU) 120C. Also provided in control unit 120 are a signal generatorutility 122 (e.g. function generator and phase control) configured andoperable to control the operation of acoustic unit (transducerarrangement) 110, and an appropriate utility 123 configured foroperating optical unit 101C. Data processing and analyzing utility 120Cis preprogrammed for receiving measured data (MD) coming from detectionassembly 102A (via cable 105 in the present example) and for processingthis measured data to identify the detected light distributioncorresponding to measurements locations in the region of interest,thereby enabling determination of one or more desired parameters of theregion of interest, e.g., oxygen saturation level. Also provided in thecontrol unit is a correlator utility 125 (typically a software utility)associated with the signal generator 122.

According to this example, measurement unit 101 is configured as a probehaving a support structure (preferably flexible) 403 to be put incontact with the body part to be measured. Support structure 403 carriesat least part of illumination assembly 101A and at least part ofdetection assembly 102A. As shown in the figure, provided on the probeare: a light output port OP (constituting a lighting element) associatedwith the illumination source, a light input port IP (constituting adetecting element) associated with the detector unit, and an acousticport 245 associated with the acoustic unit. It should be understood thatlight output port OP may be integral with the light emitting element(s)or may be constituted by the distal end of an optical fiber unitconnected at its other end to one or more light emitting element(s)located outside the support structure (e.g., at the control unit).Similarly, light input port IP may be integral to with the lightdetecting element(s) or may be constituted by the distal end of anoptical fiber unit which by its other end is connected to one or moredetecting elements (light sensors) located outside the support structure(e.g., at the control unit).

Generally, illumination assembly 101A can be configured to produce lightof at least one wavelength. According to an embodiment of the presentinvention, the illumination assembly generates light of multiple (atleast two) different wavelengths. Illumination assembly 101A may forexample be preprogrammed to produce the different wavelength componentsat different times, or to simultaneously produce wavelength componentswith different frequency- and/or phase-modulation. Accordingly, controlunit 120 is preprogrammed to identify, in a signal generated bydetection assembly 102A, the corresponding wavelength of light, usingtime, and/or phase, and/or frequency analysis. The detection assemblymay include an appropriate frequency filter.

Thus, illumination assembly 101A may include the light emitter(s)carried by support structure 403 and communicating with control unit 120(using cables 106 or wireless signal transmission). Alternatively, thelight emitter(s) may be located outside support structure 403 (e.g.,within control unit 120) and connection 106 is constituted by a lightguiding assembly (e.g., optical fibers) for guiding light to lightoutput port OP located on support structure 403. Detection assembly 102Aincludes one or more light detectors such as a photomultiplier tube,photodiode or an avalanche photodiode. The light detector may include animage pixel array, e.g., CCD or other array of photodiodes. Thedetector(s) may be accommodated outside support structure (probe) 403,e.g., may be located within control unit 120, and returned light (lightresponse) may be guided from input port IP of the detection assembly vialight guiding means 105 (e.g., optical fibers). Alternatively, thedetector(s) may be located at the support structure and connection 105is configured to connect an electrical output of the detector(s)indicative of measured data MD to control unit 120. As indicated above,detection assembly 102A may include two separate detectors or an arrayof detectors. It should also be understood that connections 105 and 106may be electric wires connecting control unit 120 to the illuminationassembly and detection assembly located on support structure, 403, orthe connection may be wireless.

Thus, generally, the terms “illumination assembly” and “detectionassembly” as carried by a support structure (probe) which is brought toa body part to be measured, are constituted by at least lighttransmitting and receiving ports. Similarly, transducer arrangement 110may be located on support structure 403 (so as to be brought in acousticcontact with the skin), and connected to control unit 120 (its signalgenerator 122 and CPU 120C) using cables and/or optical fibers 107and/or using wireless means. Alternatively, connection 107 mayconstitute an acoustic guiding unit for connecting the transducer(s)located outside the support structure (e.g., at the control unit) toacoustic output port 245 on the support structure.

Transducer arrangement 110 may be a single acoustic element, configuredand operable for emitting focused or unfocused acoustic beams or foremitting acoustic pulses; or a piezoelectric phased array capable ofproducing acoustic beams with variable direction, focus, duration andphase; or may be an array of silicon units or other pressure generatingunits configured as a single element or an array of elements (phasedarray); or a complete ultrasound imaging probe comprising transmittingand receiving units. The transducer arrangement may be connected to anamplifier (not shown), e.g. located within control unit 120, operable toamplify electronic signals generated by signal generator 122. Thecontrol unit is preprogrammed to operate transducer arrangement 110 (viasignal generator 122) in a predetermined manner to produce a codedacoustic continuous wave, which is a predetermined function of at leastone parameter of the acoustic radiation varying over time during ameasurement time interval. This predetermined function is selected tohave a narrow autocorrelation function (i.e. an autocorrelation which isnegligible for any delay time larger than the determined time resolutionof the system, for example, determined as the time resolution ofdetection of the electromagnetic radiation response, or as the temporalbandwidth of the ultrasound transducer, or as the required spatialresolution divided by the speed of sound in the media), as will bedescribed more specifically below.

Detection assembly 102A generates electronic signals in response to theamplitude and phase of light collected at input port IP. Theseelectronic signals may be filtered by analog and/or digital filters, forexample bandpass filters, that are appropriately provided beingconnected to data processing utility 120C of control unit 120 or being apart of this processing utility.

Reference is now made to FIG. 2A showing more specifically an example ofthe functional elements and operation of control unit 120. As shown,signal generator 122 to includes a signal source 124, a modulator 126and a sequence generator 128. Also, in the present example, the controlunit includes a phase shifter utility 127, as in the present examplephase is a modulatable parameter in a continuous acoustic wave to begenerated by the acoustic unit. However, the invention is not limited tothis specific example, and the control unit may include a frequencyand/or amplitude shifter utility, alternatively or additionally to thephase shifter. CPU 120C controls the operation of signal generator 122and cross correlator 125, and receives signals from cross correlator125. CPU 120C also controls signal source 124 and sequence generator128. Signal source 124 generates a base signal S₀ (e.g. a sine wave of acertain frequency, or a chirped signal, or a train of square waves at acentral frequency). Signal S₀ passes through modulator 126 that controlsone or more of its parameters (e.g. at least one of the following:phase, frequency, frequency gradient (chirp), phase jump, amplitude,duty cycle, chirp gradient) to produce a modulating (or coding) signalS₂ to operate the acoustic transducer to produce an acoustic(ultrasound) sequence. Alternatively modulating signal S₂ comprises acombination of such sequences. The operation of modulator 126 iscontrolled by sequence generator 128. CPU 120C transmits a signal S₁ tosequence generator 128, that controls its operation. Sequence generator128 in turn controls the operation of modulator 126 according to thissignal S₁. Modulating signal S₂ that exits modulator 126 is a result ofthe combination of base signal S₀ with the modulation induced bymodulator 128 on this signal. This signal S₂ is transmitted totransducer arrangement 110 via connection 107. An additional poweramplifier can be used to amplify signal S₂ before actuating thetransducer. Signal S₂ is also transmitted to cross correlator 125, whichcorrelates this signal S₂ with measured data MD coming from detectionunit 102A through connection 105. Alternatively, cross correlator 125may correlate a signal S₃ corresponding to signal S₂ with the measureddata. Corresponding signal S₃ is for example the amplitude of signal S₂or its absolute value, or another function corresponding to signal S₂.Phase shifter 127 controls the phase of signal S₃, such that a phaseshift is generated between signal S₂ and signal S₃. The output of crosscorrelator 125 (e.g. the amplitude or phase of the cross correlation atdifferent delays) is processed by CPU 120C and displayed on display120A.

Reference is made to FIG. 2B exemplifying a method of the presentinvention suitable to be used to extract the light distribution in thetissue. An arbitrary waveform (GWF) is generated with predeterminedcharacteristics such that the autocorrelation of to GWF is negligiblefor any delay z larger than the time resolution of the system. Thiswaveform GWF is saved to memory. This arbitrary waveform corresponds tothe above-described modulating signal S₂.

The GWF is transmitted to actuate an ultrasound transducer (110 inFIG. 1) with a known bandwidth, for producing acoustic waves in the formof a non-periodic sequence to irradiate a volume of medium (tissue), atleast part of a region of interest (200 in FIG. 1). Concurrently, theillumination assembly is operated to illuminate the medium with coherentlight of a certain wavelength 2. This light propagates through the samevolume through which the acoustic waves propagate (tagged volume), andlight returned from the medium is detected (representing a lightresponse of the medium).

Electronic signals generated by the detection assembly in response tothe detected light are stored in memory, using a sampling card with asampling frequency, which is at least twice the transducer's bandwidth,thereby enabling exact reconstruction of a continuous-time signal fromits samples. These signals are cross correlated against the GWFelectronic signals, or against a function of the GWF signal as describedbelow, stored in memory with different time delays as applied. For eachdelay τ, the amplitude of the cross-correlation (CCA(τ,λ)) is stored inmemory. According to a preferred embodiment of the present invention,CCA(τ,λ) represents the light distribution at wavelength λ multiplied bythe acoustic power distribution or pressure amplitude, or a function ofthe acoustic pressure amplitude (PA(τ)) at a distance z corresponding tothe product of τ and the speed of sound c_(s) in the measured tissue(i.e. z=τ·c_(s)).

Pressure profile (PA(τ)) may or may not be known. In the case where theoverall output light distribution function is a product of the pressureprofile and the light distribution function within the medium, if thepressure profile is known, the light distribution LD at wavelength λ isdetermined as

LD(z,λ)=CCA(z,λ)/PA(z)  [1]

In case the (PA(τ)) is unknown, the measurements are performed using atleast two different wavelengths λ₁, λ₂ of light providing twocorresponding cross-correlation amplitudes CCA(z, λ₁) and CCA(z, λ₂)respectively. A ratio between the two measurements is independent of(PA(τ)), thus providing the ratio between the light distributionsdetermined as

$\begin{matrix}{\frac{{LD}\left( {z,\lambda_{1}} \right)}{{LD}\left( {z,\lambda_{2}} \right)} = {\frac{{CCA}\left( {z,\lambda_{1}} \right)}{{CCA}\left( {z,\lambda_{2}} \right)}.}} & \lbrack 2\rbrack\end{matrix}$

This will be described more specifically further below.

An example of an acoustic sequence used in an embodiment of thisinvention utilizes a random number generator with a long enough period(infinite relative to the used segment). An example of such a functionmight be:

S ₂ =A ₁ cos(ωt+Θ(i))  [3]

where A₁ is the amplitude, i=floor(t/τ) and Θ is a sequence of randomnumbers in the range [0,2π] generated by any rectangular pseudo randomgenerator.

In this example, the phase of signal S₂ of angular frequency ω (producedby modulator 126 in FIG. 2A) is randomly shifted for each segment ofduration τ. Here, τ determines the width of the auto correlation of thesignal around zero, and hence the spatial resolution of the processedsignal. The lowest limit of a range of values of τ may be bound by thebandwidth of the acoustic system, which is proportional to 1/τ.

FIGS. 3A and 3B show, respectively, a segment of typical signal S₂ andits auto correlation. The tagging of incident light is detected as aresult of interaction between acoustic waves and photons whose opticalpath and consequently phase is modulated by motion of scatteringparticles in the tissue. The phase of the received signal (correspondingto the tagged photons) relative to the phase of transmitted signal S₂varies with time and depth and is also unknown.

Hence, processing of the measured data indicative of the detected lightresponse includes processing of the correlated signal to search for thephase shift that gives the best correlation for each delay. This can bedone by correlating the measured data with a complex phasorrepresentation of the acoustic sequence and taking the absolute value ofthe resulting phasor. Considering the above example for signal S₂, thecomplex phasor is S_(p)=e^(jωt+Θ(i)) and the correlation C(τ) is to becalculated as

$\begin{matrix}{{C(\tau)} = {\sum\limits_{t}{{D(t)}{S_{p}\left( {t - \tau} \right)}}}} & \lbrack 4\rbrack\end{matrix}$

where D is the measured data (MD in FIG. 1); S_(p) is an example forsignals S₃ that can be used to determine the correlation between S₂ andthe measured data.

If the signal originates from a known delay τ₀ with arbitrary phaseshift φ, the result would be:

$\begin{matrix}{{C\left( \tau_{0} \right)} = {{\sum\limits_{t}{{\cos \left( {{\omega \; t} + {\Theta (i)} + \varphi} \right)}{S_{p}\left( {t - \tau_{0}} \right)}}}}} & \lbrack 5\rbrack\end{matrix}$

FIG. 3C illustrates the output C(τ₀) for τ₀=10⁻⁵ seconds and any valueof φ.

A measurement interval, i.e. the duration of signal S₂ in onemeasurement epoch, should preferably be as long as possible to enhancethe signal to noise ratio (SNR) of the system. This can be implementedunder the assumption that a scattering pattern is constant during themeasurement, and therefore a phase relation of the measured data and theoriginal signal S₂ is constant (even if such a relation is unknown). Thetemperature related Brownian motion of the scattering particles andother effects in a live tissue, such as motion of blood cells, enforce apractical upper limit for the measurement duration. These motions causethe interference pattern between the different photons on the sensingsurface of a detector (called “speckle pattern”) to be time varying, andcause the phase relation between the measured data and signal S₂ to varywith time. The measurement interval is therefore bounded by the specklecorrelation time, defined, for example, as disclosed in Lev et al. in J.Opt. Soc. Am. A Vol. 20, No. 12 (December 2003).

Signal S₂ may comprise a sequence of short pulses, that are separated byperiods of low (or even zero) amplitude. The separation period betweenpulses is determined as the time period where the phase of light thatpropagates through the media during the second pulse is independent ofthe phase of light that propagates through the media during the firstpulse. Preferably, the separation time should be longer than the specklecorrelation time. As the speckle correlation time depends on theproperties of the media (such as its temperature), the signal S₂ can bedetermined according to the properties of the media being monitored. Byanother option, a plurality of separation durations can be used, and anoptimal separation duration (providing optimal SNR or an optimizedsignal parameter) should be selected for measurement. According to yetanother option, the optimal separation itself may be monitored, toprovide a measure for a property of the media (such as its temperatureor the flow of blood through the tissue).

In some cases, the above pulses can be replaced each by a set ofmultiple pulses that are transmitted over the same phase of light—thusto allow obtaining strong enough signal by means of averaging. Theseparation durations between these inner-set of to pulses is selected tobe large enough so that during the propagation of a single pulse(including echoes) through the region of interest the pulses do notco-exist inside the region of interest, and to be smaller than thespeckle correlation time.

If longer integration is required to further improve the SNR, averagingcan be carried out between separate measurements' intervals, but thisaveraging is done after the absolute value of the complex correlation iscalculated separately for each of the measurements. In the case where S₂comprises a series of pulses, the averaging may be performed over theabsolute value of the cross correlation for each pulse separately. Forexample, averaging can be performed over a predetermined number ofmeasurements that are separated by a predetermined time delay. Thisaveraging might be advantageous in cases where the measured data isperiodic (i.e. changes periodically as a function of time as in the caseof modulation of the blood volume). For example, averaging overdifferent portions of measured data can be correlated with thepeaks/troughs of the blood volume during systolic/diastolic periods in apulsating blood volume, having a predetermined delay from each other. Inthis case, a difference between the signals corresponds to the oxygensaturation levels of blood (as in the case of pulse oximetry).

Signal S₂ can for example include a plurality of different arbitrarysignals. These may for example be different signals having differentamplitudes and/or different frequencies and/or different phasevariations.

The above example demonstrates random modulation of the phase of signalS₂. As indicated above, other parameters of signal S₂ may be modulatedaccording to a predetermined function.

The following is an example related to a frequency modulation of signalS₂. Signal S₂ can be expressed by

S ₂=cos(φ(t)),  [6]

where φ is selected so that

${\frac{\varphi}{t} = \omega_{i}},$

and where i=floor(t/τ) and ω is a random sequence with a squaredistribution in the assigned angular frequency range.

FIGS. 4A and 4B exemplify, respectively, a segment of signal S₂ and itsautocorrelation. FIG. 4C shows the cross correlation C(τ) definedsimilarly to the above example for phase modulation.

It should be noted that the above description for cross correlation isbased on digital signal processing. However, dedicated analog circuitsthat perform cross correlation with variable delays can be designed andconstructed to provide a similar to functional operation of the system.

Referring to FIG. 5, there is exemplified a case for the phasemodulation of signal S₂. The figure shows the amplitude of C(τ) (i.e.amplitude of cross correlation CCA(τ,λ), obtained for different valuesof delay τ, as a function of distance from the acoustic transducer,where this distance equals to the product of τ by the speed of sound inthe medium. Three graphs are presented, showing CCA(τ,λ) calculated fromexperimentally obtained measured data MD corresponding to a lightresponse at three different wavelengths λ¹, λ², λ³, respectively.

In this example, three different light sources, at three differentwavelengths, illuminate a turbid medium, and a detection unit generateselectronic signals indicative of measured data corresponding to lightcollected at the input port of the detector, for each wavelength used.As can be seen in the figure, the amplitudes of cross correlationsignals CCA(τ,λ¹), CCA(τ,λ²), CCA(τ,λ³), or generally CCA(τ,λ^(i)), atvarying distances is different for the three wavelengths. This resultsfrom the fact that the light distribution of the three wavelengths inthe tissue is different, due to differences in absorption, scatteringand index of refraction.

Signal CCA(τ,λ^(i)) corresponds to the acoustic distribution or pressureamplitude PA(z), and to the light distribution LD(λ^(i)).

$\begin{matrix}{{{LD}\left( {z,\lambda^{i}} \right)} = {K*{\prod\limits_{{\alpha = s},d}{\left( {1 + \frac{1}{\mu^{i}\sqrt{\left( {\overset{\rightarrow}{r} - {\overset{\rightarrow}{r}}_{\alpha}} \right)^{2} + z^{2}}}} \right)\; \frac{z}{\left( {\overset{\rightarrow}{r} - {\overset{\rightarrow}{r}}_{\alpha}} \right)^{2} + z^{2}}{\exp \left( {{- \mu^{i}}\sqrt{\left( {\overset{\rightarrow}{r} - {\overset{\rightarrow}{r}}_{\alpha}} \right)^{2} + z^{2}}} \right)}}}}} & (7)\end{matrix}$

where K is a constant, μ^(i)=√{square root over (3μ_(a) ^(i)(μ_(a)^(i)+_(s) ^(i)))}≅√{square root over (3μ_(a) ^(i)μ_(s) ^(i))} is theeffective decay rate of light in the medium, μ_(a) ^(i) is theabsorption coefficient and μ_(s) ^(i) is the scattering coefficient atwavelength λ^(i); when near infrared light is used, it can be assumedthat μ^(i)≅√{square root over (3μ_(a) ^(i)μ_(s) ^(i))}, {right arrowover (r)}_(α) is either the vector to the source (α=s) or to thedetector (α=d), and z is the direction parallel to the direction ofpropagation of the acoustic radiation into the medium.

For example, for a large enough distance z (z=τ·c_(s), c_(s) being thespeed of sound in the medium) from the body surface (namely larger thanthe mean free path of light in the medium, and larger than the sourcedetector separation, {right arrow over (r)}_(d)−{right arrow over(r)}_(s), the light distribution LD(z,λ^(i)) is proportional to e^(−2μ)^(i) ^(z) where CCA(z,λ^(i))) is given by CCA(z,λ^(i))≅PA(z)I₀^(i)e^(−2μ) ^(i) ^(z)+C_(o), where I₀ ^(i) is the initial lightintensity upon entry into the medium, and C_(o) is an additive constant.

Thus, turning back to FIG. 2B, if the acoustic pressure amplitude PA(z)is known, for example by measuring it with a hydrophone in water, thelight distribution LD(z,λ^(i)) can be extracted by dividing PA(z) out ofCCA(z,λ^(i)), after eliminating C_(o). In many practical cases, however,the pressure profile is unknown, for example when the medium consists ofdifferent layers with different acoustic impedances. Thus, there is nocorrespondence between measurements of the pressure profile in water orsynthetic phantoms and the correct pressure profile in the measuredmedium. In such cases, measurements with at least two or generally Ndifferent light wavelengths can be performed, and correspondingCCA(z,λ^(i)) are used to eliminate the acoustic contribution PA(z)(after eliminating C_(o)). This is implemented by dividing measuredCCA(z,λ^(i)) by measured CCA(z,λ^(i)) for i≠j, assuming that theacoustic contribution is the same for all wavelengths, which is ajustifiable assumption. Thus, the ratio of the light distributions canbe obtained. This ratio is important for example for determining theoxygen saturation of a tissue or blood vessel as will be explainedbelow.

Constant C_(o) corresponds to the noise level of the system at themeasured frequency bandwidth. For example, one possible way to measureC_(o), is to cross correlate measured data MD with a time-reversedsignal S_(p)(τ−t). Such a correlation results in the same frequencybandwidth, but is completely uncorrelated with measured data MD. Thus,constant C_(o) for each wavelength of light can be measuredindependently and eliminated from signal CCA(z,λ^(i)). Alternatively,C_(o) can be eliminated by performing the measurements at two differentamplitudes of acoustic radiation, and taking the difference between thetwo corresponding cross correlations.

In the case of a medium irradiated by three different wavelengths:

$\begin{matrix}{\frac{{\overset{\sim}{I}}^{i}}{{\overset{\sim}{I}}^{j}} = {\frac{I_{0}^{i}}{I_{0}^{j}}^{{- 2}\Delta \; \mu^{ij}z}}} & \lbrack 8\rbrack\end{matrix}$

where i; j=1; 2; 3 represent the three lasers, Ĩ=(CCA(z,λ^(i))−C_(o)) isthe amplitude of the signal at distance z, I₀ ^(i), I₀ ^(j) are theinput intensities of the i^(th) and j^(th) wavelengths respectively andΔμ^(ij)=μ^(i)−μ^(j).

Taking a logarithm of the equation above, Δμ^(ij) can be obtained:

$\begin{matrix}{{\Delta \; \mu^{ij}} = {{- \frac{1}{2}}\frac{\partial}{\partial z}{\ln \left\lbrack \frac{{\overset{\sim}{I}}^{i}}{{\overset{\sim}{I}}^{j}} \right\rbrack}}} & \lbrack 9\rbrack\end{matrix}$

The saturation s is related to the absorption coefficient μ_(a) ^(i) bythe following relation:

μ_(a) ^(i)=ε_(Hb) ^(i) C _(Hb)+ε_(HbO) ^(i) C _(HbO) =C _(tot)(ε_(Hb)^(i)+(ε_(HbO) ^(i)−ε_(Hb) ^(i))s)  [10]

where C_(Hb), C_(HbO) and C_(tot) are the concentrations of deoxygenatedhemoglobin, oxygenated hemoglobin and the total hemoglobin,respectively, s is the oxygen saturation defined as the ratio betweenthe concentration of oxygenated hemoglobin to the total hemoglobinconcentration (i.e. s=C_(HbO)/C_(tot)), and ε^(i) _(Hb), ε^(i) _(HbO)are the extinction coefficients at the i^(th) wavelength fordeoxygenated and oxygenated hemoglobin, respectively, that are known inthe literature.

Thus, for any saturation s, the theoretical μ_(a) ^(i) can be calculatedusing this equation, in order to determine the saturation at differenttissue layers. The decay coefficient μ^(i) can be calculated for examplefrom the graphs presented in FIG. 5. If the scattering coefficient isassumed to be the same for all three wavelengths, the absorptioncoefficient at each wavelength equals:

${\mu_{a}^{i} = \frac{\left( \mu^{i} \right)^{2}}{3\mu_{s\;}}},$

and the ratio α^(ijk) can be calculated by:

$\begin{matrix}{{\alpha^{ijk} = {\frac{\Delta \; \mu^{ij}}{\Delta \; \mu^{ik}} = \sqrt{\frac{ɛ_{Hb}^{i} - ɛ_{Hb}^{j} + {\left( {ɛ_{HbO}^{i} - ɛ_{HbO}^{i} - ɛ_{Hb}^{i} + ɛ_{Hb}^{j}} \right)s}}{ɛ_{Hb}^{i} - ɛ_{Hb}^{k} + {\left( {ɛ_{HbO}^{i} - ɛ_{HbO}^{k} - ɛ_{Hb}^{i} + ɛ_{Hb}^{k}} \right)s}}}}},} & \lbrack 11\rbrack\end{matrix}$

The following explains how the saturation is calculated by usingmeasured data obtained by three lasers, and using the differences inΔμ¹² and Δμ³¹:

The extinction coefficients are known from the literature, so that forSat=1-100% the theoretical values for Δμ_(th)=(Δμ_(th) ¹², Δμ_(th) ³¹)can be calculated up to the multiplicative constant √{square root over(3·μ_(s) ^(i)·C_(tot))}. The scattering coefficient μ_(s) ^(i)=μ_(s) isapproximated to be the same for the three lasers, however it may varywith time. In order to compare the experimental value Δμ_(ex)=(Δμ¹²_(ex),Δμ³¹ _(ex)) to the theoretical value Δμ_(th), the angle betweenthe vectors in the plane that is spanned by [Δμ¹²,Δμ³¹] is determinedFor each experimental point there is a certain value of Δμ_(ex). Theangle between this experimental vector and every theoretical option(corresponding to saturation values of 1%-100%) is calculated. Thesaturation value that corresponds to Δμ_(th), which has the smallestangle to Δμ_(ex), is the calculated saturation level. Thus, thesaturation is calculated without depending on the factor √{square rootover (3·μ_(s) ^(i)·C_(tot))}.

Experimental data of the graphs presented in FIG. 5 was collected whenthe distance between the illumination and detection units was 3 cm. Thepeak intensity was obtained at about 9 mm from the skin (there is a 2-3mm distance between the transducer face and the skin surface in thismeasurement). In order to map the three dimensional light distribution,different separations between the source and the detectors should beused.

Once the saturation s is determined, the total hemoglobin concentrationC_(tot) can be determined from measurements of the exponential decay ofCCA(z, λ^(i)) at the different wavelengths, using the known extinctioncoefficients for oxygenated and deoxygenated hemoglobin.

It should be noted that, in addition to the oxygen saturation level,other parameters of the tissue and blood composition or parameters canbe determined from measurements of CCA(z, λ^(i)). Moreover, the presentinvention provides for using determination of Δμ_(ex), without relyingon measuring CCA(z, λ^(i)), for example by using frequency domainspectroscopy or time of flight based measurements, to determine thefollowing parameters independent of the measurement method:

For example, total Hemoglobin content C_(tot) can be calculated asfollows: Since the angle between the vectors Δμ_(th) and Δμ_(ex)corresponds to the calculated saturation, the multiplicative factors(i.e. √{square root over (3·μ_(s) ^(i)·C_(tot))}) that are neglected inthe theoretical calculation of Δμ_(th) are of no consequence to thesaturation value that results from the disclosed algorithm. If there isa change in the total blood concentration, C_(tot), or the scatteringcoefficient, μ_(s), without changes in the oxygen saturation level, itwill be reflected by the distance of the experimental point from theorigin (see Eq. [10]), but the direction of the vector from the originto the experimental point will remain the same. Therefore, the totalblood concentration can be measured by determining the distance of theexperimental Δμ_(ex) point from the origin. Changes in the scatteringcoefficient can be extracted using other optical methods, such as timeof flight or frequency domain spectroscopy. Consequently, independentmeasurements of the total blood concentration C_(tot) and the scatteringcoefficient μ_(s) can be made.

Another parameter that can be determined from measurements of CCA(z,λ^(i)) is blood flow. In general, the measured tissue volume containsblood vessels and capillaries. The flow of blood inside these vesselsaffects the properties of the measured data. The speckle correlationtime is affected by the flow, there is a flow dependent Doppler shift inthe acoustic waves and other effects may exist. Direct measurement ofthe speckle correlation time is known to correspond to blood flowvelocities [G. Yu et al Journal of Biomedical Optics 2005 10:2]. Thus,the properties of CCA(z, λ^(i)), such as the peak amplitude, the noiselevel Co and other parameters are affected by the flow. By monitoringthese parameters, as a function of time, changes in the flow rates areextracted. In particular, by monitoring these changes as a function ofdepth, flow distribution can be determined

Yet other measurable parameters include differences between arterial andvenous contribution to the signal. In this connection, the followingshould be noted: General Near Infrared Spectroscopy (NIRS) measurementsdo not distinguish between the arterial, capillary, and venouscompartments of blood circulation and thus reflect a weighted average ofHb concentrations within these different blood compartments in theregion sampled. For example, in brain, the relative distribution ofarterial, capillary, and venous compartments in the cerebral bloodvolume (CBV) is generally accepted to be approximately 20%, 10%, and 70%respectively. Using this distribution, C_(tot) in the venous compartmentcan be isolated as follows:

C _(tot)=0.2[Hb]a+0.1[Hb]c+0.7[Hb]v

where C_(tot), [Hb]a, [Hb]c, and [Hb]v are the concentrations of totalHb, arterial Hb, capillary Hb, and venous Hb, respectively. Using theassumption that the capillary concentration of Hb is the mean ofarterial and venous concentrations, it is possible to determine [Hb]v,given that C_(tot) can be measured and [Hb]a can be calculated from thearterial saturation SaO₂ using measured Hb content of arterial blood andCBV as a measure of the percentage of blood in a given tissue volume.SaO2 can be measured using a pulse oximeter. Because Hb is generated inthe brain solely through the process of O₂ dissociation from HbO₂, thedifference in [Hb]a and [Hb]v is identical, although opposite in sign,to the difference in [HbO₂]a and [HbO₂]v, assuming that CBV remainsconstant during the measurement period.

Yet another parameter that can be determined, based on measurements ofCCA(z, λ^(i)), is the oxygen extraction fraction (OEF). OEF is thepercentage of oxygen extracted from arterial blood in the tissue:

OEF_=(arterio-venous O₂ diff)/CaO₂,  [12]

where CaO₂, the arterial oxygen content, can be calculated from thearterial saturation (measured by a pulse oximeter for example, asexplained by Brown D. W. et al. Pediatric Research Vol 54 No 6 2003 pp861-867); and

(arterio-venous O₂ diff)=([Hb]v _(—) −[Hb]a)*1.39 ml O₂/gHb  [13]

where [Hb]v and [Hb]a are defined above.

Therefore, as the total hemoglobin content C_(tot) can be extracted asexplained above, the oxygen extraction fraction in the measured tissuevolume can be determined using the preferred embodiment.

Reference is now made to FIGS. 6A-6C schematically illustrating threeexamples, respectively, of the probe configurations according to furtherembodiments of the invention. In these examples the probe includes anannular acoustic transducer unit 110 and light input and output ports IPand OP such that at least one of these ports is located within anannular aperture of the acoustic transducer module. In the example ofFIG. 6A, a common light guiding unit 310 is used through which fibers105 and 106, associated with the light output and input ports OP and IP(i.e. with the lighting and detecting elements), pass. In the example ofFIG. 6B, the configuration is generally similar to that of FIG. 6A, bututilizes separate light guiding units 312 and 311 located inside thetransducer unit's aperture and associated with lighting and detectingelements, respectively. In the example of FIG. 6C, a light guiding unit310 carrying an optical fiber 306 associated with light detectingelement IP is located inside the transducer's aperture, and a lightingelement OP is located outside the transducer unit adjacent thereto beingconnected to a light source or control unit via an appropriateconnection 105. The location of elements corresponding to lightingelements and light detecting to elements can be interchanged.

FIGS. 7A and 7B schematically illustrate two more examples,respectively, of a probe configuration according to the invention. Inboth of these examples, an acoustic transducer unit 110 is configuredfor passage therethrough of at least one light guiding unit associatedwith light detecting element(s), and an illumination unit includes aplurality of lighting elements located outside the transducerarrangement adjacent thereto. The lighting elements are arranged in acircular array around the light detecting element(s). In the example ofFIG. 7A, a single light detecting element IP is used being located in anaperture of the transducer and associated with an appropriate lightguiding unit 310 (e.g. fiber), and a circular array of twelve lightingunits 320A-320L is used. In the example of FIG. 7B, an illumination unitincludes eight light detecting elements are used located incorresponding spaced-apart apertures of the transducer. The lightdetecting elements include a central element 310A and elements 310Barranged in a circular array around element 310A. The location ofelements corresponding to lighting elements and light detecting elementscan be interchanged.

Reference is made to FIGS. 8A and 8B exemplifying the transducer'sassemblies including light guides.

FIG. 8A shows a transducer's assembly 500 that includes a light guidingelement in the center. The assembly may be configured such that there isno acoustic contact between the light guiding element (an optical fiber)and the piezoelectric element that generates the acoustic waves.Assembly 500 comprises a casing 501 that encapsulates piezoelectricelement 502 and light guide 510. To allow for light delivery throughpiezoelectric element 502, the piezoelectric element 502 is formed withan optical window 509 having a diameter large enough to provide lightpropagation from/to light guide 510 through this optical window 509.Optical window 509 may be a physical hole, or may be a transparentopening in the piezoelectric element. Piezoelectric element 502 may alsobe completely transparent to light, and therefore optical window 509 maybe a part of piezoelectric element 502. In addition, the optical window509 may include a transparent optical rod that will allow lightpropagation through. In case the optical window 509 is a physical hole,an additional optical window 503 can also be used to seal this hole 509.Further provided is a support 511 configured to allow aligning of theoptical window 509 with the aperture of light guide 510. Electric wires505 and 506 are coupled to the two electrodes (not shown) on thepiezoelectric element 502, for to generating acoustic waves. These twowires are connected to cable 107 (see FIG. 1) used to deliver electricalsignals from the signal generator (125 in FIG. 1).

FIG. 8B shows an assembly 551 including a casing 501 that encapsulatespiezoelectric element 502 and optical guide 510. Optical guide 510enters the casing 501 through an opening (not shown) and is supported bya support structure 515 inside the casing. Another support structure556, positioned inside the casing, supports a prism 575. The prism ispositioned such that light coupled from optical guide 510 is directedtowards a further optical guide 565. This optical guide 565 ispositioned inside a through hole (optical window) 509. The piezoelectricelement 502 is supported by a support structure 555 that preventsacoustic coupling to the casing walls. Electric wires 505 and 506 arecoupled to the two electrodes (not shown) on the piezoelectric element502, for generating acoustic waves. Alternatively, the optical guide 510may be input from the side its end cut at an angle to allow for thelight trapped inside the fiber to reflect at a 90° angle, i.e. a sidefiring fiber, instead of propagating through the prism 575.

The above configurations allow for selecting the light input and outputports for use in measurements so as to provide an optimal distancebetween the operative input and output ports. This is associated withthe following: As the distance between the light source and lightdetector is reduced (to ˜zero), the contribution of light reflected fromsuperficial layers to the untagged signal in the detected light ishigher than in the case of larger source-detector distance. Therefore,in order to detect the tagged light from deep layers, the detection unitpreferably includes an electronic filter, one of the kind that filtersthe low frequency signals generated in response to untagged light fromthe signals corresponding to tagged light (at higher frequencycorresponding to the ultrasound bandwidth). Reducing the source-detectordistance also improves the accuracy in calculating the opticalproperties of the medium, improving the determination of the desiredparameter(s), e.g. calculation of the oxygen saturation level. When thesource-detector distance is small, the differences in the optical pathsof the shallow photons and the deep photons (that are used to calculatethe optical attenuation coefficient) depend primarily on the distancetraveled in the z direction (along the radiation direction towards theregion of interest). Whereas for larger source-detector distance, theoptical attenuation also depends on the differences in the x and ydimensions, and thus degrades the dependence on the z direction,rendering the to calculations more complex.

As a result of the ultrasound beam interacting with the light, thesignal that we obtain includes an integral over {right arrow over (r)}of LD(λ^(i)) within V_(US) the volume of the ultrasound beam

$I = {\int_{V_{US}}{{\overset{\rightarrow}{r}}{{{LD}\left( \lambda^{i} \right)}.}}}$

This integral will clearly depend on {right arrow over (r)}_(s) and{right arrow over (r)}_(d). The expression for the light distributionLD(z) (Eq [7]) shows that its integral over r depends on thesource-detector distance r_(sd), such that as distance r_(sd) decreasesthe light distribution LD(z) will depend primarily on the exponentialdecay.

In addition, at large source-detector distances, there are many morescattering events of photons reaching the detector than for smallsource-detector distances. Thus, as different wavelengths are scattereddifferently by the tissue and cells, the difference between the opticalpaths of the different wavelengths increases as the source-detectordistance increases. Since it was assumed above that the scatteringcoefficient is the same, the error in making this assumption increasesas the source-detector distance increases.

Thus, the present invention provides for an effective technique fordetermining one or more desired parameters of a subject using anacoustic tagging of light, where acoustic radiation is generated in theform of a continuous wave, which is coded (modulated) to vary inaccordance with a predetermined function of at least one parameter ofthe acoustic radiation which is non-periodic over a measurement timeinterval. The invention also provides an optimized probe configurationto obtain a required distance between the light input and output portsused in the measurements.

Those skilled in the art will readily appreciate that variousmodifications and changes can be applied to the embodiments of theinvention as hereinbefore described without departing from its scopedefined in and by the appended claims. In the method claims that follow,alphabetic characters and numerals used to designate claim steps areprovided for convenience only and do not imply any particular order ofperforming the steps.

1. A probe device for use in a system for non invasively monitoring atleast one parameter of a region of interest in a subject's body, theprobe device comprising: a support structure configured to be put incontact with the body part, the support structure comprising: anarrangement of light output, light input and acoustic output portsassociated with respectively an illumination assembly, a detectionassembly, and an acoustic unit configured and operable for illuminatinga region in the subject's body by light coming from at least one lightoutput port, irradiating a region in the body with acoustic waves tocoming from at least one acoustic output port, and collecting lightscattered from the illuminated region in the body by at least one lightinput port; and a control unit configured and operable to select atleast one light output port, at least one light input port, and at leastone acoustic output port as operative ports to provide an operatingcondition of the probe during a measurement session, the operatingcondition corresponding to a desirably short distance between theoperative light output and light input ports, such that the acousticwaves of a predetermined frequency range coming from said at least oneoperative acoustic output port and illuminating light coming from saidat least one operative light output port overlap in a region within aregion of interest in the body, thereby inducing tagging of lightscattered from the overlapping region by said acoustic waves andcollecting the tagged light by said at least one selected operativelight input port.